Scintillation crystal radiation detection systems rely on high-energy photons, such as gamma rays, interacting with a scintillation material in a Compton scattering or photoelectric interaction. The scintillation event produces a large number of lower-energy photons that are more readily detected using a photodetector, for example, a photomultiplier tube, silicon photomultiplier, or the like.
Exemplary scintillation crystals include NaI(TI) (thallium-doped sodium iodide), BGO (bismuth germinate), LSO (lutetium oxyorthosilicate), GSO (gadolinium orthosilicate), LYSO (cerium-doped lutetium yttrium orthosilicate), LuAP (lutetium aluminum perovskite), LGSO (Lu0.4Gd1.6SiO5: 22.0 mol % Ce), LaBr3 (lanthanum bromide) and the like. For example, when BGO interacts with high-energy radiation, such as gamma-rays or x-rays, it emits a green fluorescent light with a peak wavelength of 480 nm. BGO is used for a wide range of applications in high-energy physics, nuclear physics, space physics, nuclear medicine, geological prospecting, and other industries. LYSO crystal has the advantages of high light output and density, quick decay time, excellent energy resolution, and moderate cost. These properties make LYSO a good candidate for a range of detection applications in nuclear physics and nuclear medicine, which require improved timing and energy resolution.
In typical scintillation crystals used in positron emission tomography (PET), for example, an incident gamma photon having a nominal energy of 511 keV interacts in the scintillation crystal to produce tens of thousands of low-energy (e.g., visible wavelength) photons (˜1 eV) in a very short flash or scintillation event. The number of scintillation photons produced in the crystal is proportional to the energy deposited by the photon.
The lower-energy scintillation photons are then detected with photodetectors that are typically placed in an array on one side, or on opposite sides of the scintillation crystal. Typical photodetectors include photomultiplier tubes (PMT), avalanche photodiodes (APDs), Si-PIN photodiodes, silicon drift photodiodes, and silicon photomultipliers (SiPM). The radiation detector is configured to identify a high-energy photon interaction by detecting the low-energy photons produced in the scintillation event, and to determine the location of the scintillation event within the scintillation crystal (preferably, in three spatial dimensions), the time of the scintillation event, and the total energy of the event.
Positron Emission Tomography (PET)
Although radiation detectors in accordance with the present invention are contemplated to have applications in several different fields, ranging from cosmological imaging to the detection of radioactive materials, a particular application in the field of medical imaging will be described in some detail to assist the reader in understanding the description of the invention that will follow.
Positron emission tomography (PET) is a medical imaging modality that takes advantage of radioactive decay to measure metabolic activities inside a living organism. A PET imaging system comprises three main components, indicated schematically in FIG. 1, a radioactive tracer that is administered to the subject to be scanned, a scanner that is operable to detect the location of radioactive tracer (indirectly as discussed below), and a tomographic image processing system.
The first step is to produce and administer a radioactive tracer 90, comprising a radioactive isotope and a metabolically active molecule. The tracer 90 is injected into the body to be scanned 91. After allowing time for the tracer 90 to concentrate in certain tissues, the body 91 is suitably positioned inside the scanner 92. The radioactive decay event for tracers used in PET studies is positron emission. An emitted positron travels a short distance in the body tissue until it interacts with an electron. The positron-electron interaction is an annihilation event that produces two 511 KeV anti-parallel photons. The scanner 92 is adapted to detect the pair of photons from the annihilation event simultaneously.
The scanner 92, the second component of PET system, includes a ring of sensors that detect the 511 KeV photons, and front-end electronics that process the signals generated by the sensors. As discussed above, the scintillators 93 converts the 511 KeV high-energy photons into many lower-energy photons, typically visible light photons. The photodetectors (e.g., PMT, SiMP or APD) 94 detect the visible light photons and generate a corresponding electrical pulse. The pulses are processed by front-end electronics to determine the parameters or characteristics of the pulse (e.g., energy, timing). The front-end electronics may include, for example, one or more low-pass filters 96, analog-to-digital converters 97, and field programmable gate arrays 98. The sensors typically comprise scintillators 93 and photodetectors 94.
The data is sent from the front-end electronics to a host computer 95 that performs tomographic image reconstruction to turn the data into a 3-D image.
A 511 KeV photon has a substantial amount of energy and will pass through many materials, including body tissue. While this typically allows the photon to travel through and exit the body, the high-energy photons are difficult to detect. Photon detection is the task of the scintillator 93. The scintillator 93 absorbs or otherwise interacts with high-energy photons and emits a relatively large number of lower-energy photons, typically visible light photons. The scintillator 93 can be made from various materials, including plastics, organic and inorganic crystals, and organic liquids. Each type of scintillator has a different density, index of refraction, timing characteristics, and wavelength of maximum emission. For convenience, in the present application scintillators will sometimes be referred to as “crystals,” although other suitable scintillator materials are also contemplated. For purposes of this application “scintillator crystals” is defined to encompass any suitable scintillation material.
In general, the density of the scintillator determines how well the material stops the high-energy photons. The index of refraction of the scintillator crystal and the wavelength of the emitted light affect how easily light can be collected from the crystal. The wavelength of the emitted light also needs to be matched with the device that will turn the light into an electrical pulse (e.g., the PMT) in order to optimize the efficiency. The scintillator timing characteristics determine how long it takes the visible light to reach its maximum output (rise time) and how long it takes to decay (decay time). The rise and decay times are important because the longer the sum of these two times, the lower the number of events a detector can handle in a given period, and thus the longer the scan will take to get the same number of counts. Also, the longer the timing characteristics, the greater the likelihood that two events will overlap (pile-up), which can result in lost data.
Attached to the scintillator 93 are electronic devices that convert the visible light photons from the scintillator 93 into electronic pulses. A PMT is a vacuum tube with a photocathode, dynodes, and an anode that has high gains to allow very low levels of light to be detected. APDs are a semiconductor version of the PMT, but with significantly lower gain characteristics. SiPMs comprise an array of semiconducting photodiodes that operate in Geiger mode so that when a photon interacts and generates a carrier, a short pulse of current is generated. In an exemplary SiPM, the array of photodiodes comprises about 103 diodes per mm2. All of the diodes are connected to a common silicon substrate so the output of the array is a sum of the output of all of the diodes. The output can therefore range from a minimum wherein one photodiode fires to a maximum wherein all of the photodiodes fire. This gives theses devices a linear output even though they are made up of digital devices.
In conventional PET detectors the scintillators 93 comprise discrete crystals arranged in a two-dimensional planar array, and then arranged into a ring as shown in FIG. 1. The photodetectors 94 for detecting the flashes of scintillation light are typically positioned adjacent the back surface of each individual crystal. The signals from the photodetectors are analyzed to estimate the x-y location of the scintillation event, to estimate the depth of interaction (i.e., z-location), to determine the time of the interaction, and to estimate the total energy deposited in the scintillator. However, given the small crystal cross-sections required to obtain very high resolution, discrete crystal designs are typically expensive, have low packing fraction, reduced light collection, and are labor intensive to build.
The present inventors have researched and developed new and advanced detectors for PET scanners. For example, cMiCE: a high resolution animal PET using continuous LSO with a statistics based positioning scheme, J. Joung, R. S. Miyaoka, T. K. Lewellen, Nuclear Instruments & Methods in Physics Research A 489 (2002) 584-598 (Elsevier), which is hereby incorporated by reference in its entirety, a continuous miniature crystal element (cMiCE) detector for small animal scanners is discussed. See also, U.S. Patent Application Publication No. 2010/0044571, which is also hereby incorporated by reference in its entirety.
In another example, New Directions for dMiCE—Depth-of-Interaction Detector Design for PET Scanners, T. K. Lewellen et al., IEEE Nucl Sci Symp Conf Rec (2007); 5:3798-3802, which is hereby incorporated by reference in its entirety, a novel depth-of-interaction (DOI) detector design based on light sharing between pairs or quadlets of crystals is discussed. See also, PCT Application Publication No. WO 2010/048363, which is also hereby incorporated by reference in its entirety.